Radio-frequency coil and method for resonance imaging/analysis

ABSTRACT

An RF coil which includes an RF coil primary, an RF coil secondary, and coupling impedances that connect the two at several points along the coil peripheries. The coupling impedances are primarily reactive, and may be fixed or variable, depending on the purpose of use.

TECHNICAL FIELD

The present invention relates to resonance systems such as magneticresonance imaging (MRI) systems, and more particularly to aradio-frequency (RF) coil and method for use in such systems.

BACKGROUND OF THE INVENTION

In MRI systems or nuclear magnetic resonance (NMR) systems, a staticmagnetic field B_(o) is applied to the body under investigation todefine an equilibrium axis of magnetic alignment in the region of thebody under examination. An RF field is then applied in the region beingexamined in a direction orthogonal to the static B_(o) field direction,to excite magnetic resonance in the region, and resulting RF signals aredetected and processed. Generally, the resulting RF signals are detectedby RF coil arrangements placed adjacent to the body. See for example,U.S. Pat. No. 4,411,270 to Damadian and U.S. Pat. No. 4,793,356 to Misicet al. Typically such RF coils are either surface type coils or volumetype coils, depending on the particular application. Normally separateRF coils are used for excitation and detection, but the same coil or anarray of coils may be used for both purposes. For multiple surface RFcoils for use in NMR, see U.S. Pat. No. 4,825,162 to Roemer, et al., andfor multiple volume RF coils for use in NMR, see U.S. Pat. No. 5,258,717to Misic, et al.

A surface or volume RF coil is frequently used in examining the anatomyunder investigation to obtain, for example, an image or spectroscopic orvascular data of the right symptomatic knee. The RF coil is placed veryclose to the symptomatic knee, adjacent the region to be imaged.Unfortunately, conventional RF coil arrangements suffer from thedisadvantage that the RF coil picks up noise signals from around theregion of interest (e.g., from the contralateral asymptomatic knee).This often results in a reduced overall signal-to-noise ratio (S/N) fromthe region of interest. Similar problems occur when imaging the breasts.Yet another example is when imaging the torso, with the arms close tothe torso, or vice versa. Such problems are due in part to the fact thatthe range of the RF coil (i.e., the field of view (FOV) of the RF coil)covers a larger volume than the region of interest. In other words, thecoil FOV is larger than the desired FOV. This is because the B fieldprofile of the typical RF coil changes gradually in all directions dueto the magnetic field properties.

Attempts have been made to alleviate such problems by way of fluxrejection. See U.S. Pat. No. 5,382,903 to Young, and Burl and Young,Magnetic Resonance of Medicine, 36:326-330, 1996. The Young patentdescribes a mutually coupled resonant loop used to "buck" or reject thefield generated by the receiving RF coil to a minimum outside theimaging FOV. Although the extent of the mutual coupling could be alteredto a certain extent by changing the resonance of the flux rejectingloop, the induced currents depend primarily on the mutual couplingbetween coils which is a function of the physical separation betweencoils. The closer the respective coils the greater the coupling; and thegreater induced currents resulted in improved RF screening. However, theopposite induced currents in the flux rejecting loop reduced the netsignal intensity drastically over the imaging FOV of the coil. Thisextent varied with different coil spacings, which also affected thescreening efficiency.

For example, FIG. 1a is a loop type coil 20 commonly used for severalNMR applications. The coil 20 is resonated at the NMR frequency withthree similar capacitors with values C1 and two capacitors with values2C1. The rectangular loop serves as an inductance and resonates with acapacitance of approximately C1/4 at the NMR frequency. The center pointbetween the two 2C1 capacitors is a virtual ground point VG and isforced to the real ground. This prevents any currents on the shields ofthe coaxial cables and therefore obviates the need for cable traps. Thevoltage across one of the 2C1 capacitors is matched to 50 ohms using areactive network and fed to a preamplifier prior to system amplificationand digitization (not shown).

The square box across the C1 capacitor represents a decoupling network22 used for decoupling the RF coil during RF transmit. The number ofdecoupling networks 22 used in the RF coil 20 depends on the size of theRF coil and the degree of decoupling needed to minimize image artifactsand allow for safe preamplification operation. Additional circuitry maybe used between the RF coil 20 and the preamplifier to reduce the amountof RF energy seen at the input of the preamplifiers during whole body RFtransmit.

As shown in FIG. 1b, the decoupling network 22 consists of an inductorL1 and a pin diode D1. The pin diode D1 can be forward or reverse biasedusing programmable DC signals from the MRI system. Normally the pindiodes D1 will be forward biased such that the L1-C1 circuit is resonantat the NMR frequency. This will create a parallel trap, which willpresent an open circuit with respect to the transmitting RF currents.During receive the pin diode D1 is reverse biased. Therefore, the L1-D1circuit will effectively be open, and the RF coil will be resonant (withC1) at the NMR frequency.

FIG. 2 illustrates an RF loop coil 30 as described in the aforementionedBurl and Young article. Here, a secondary loop 34 is placed on one sideof an RF coil primary 32. Both loops 32 and 34 are magnetically coupledthrough space, and hence are mutually coupled to one another. Theprimary loop 32 is resonant with capacitance C2 and the secondary loop34 is resonant with a capacitor C3. Together the coil 30 is tuned to theNMR frequency. The capacitors in the primary and secondary loops areparalleled and include similar decoupling networks 22 as in FIG. 1.Although the mutual coupling between the primary and secondary loops canbe varied to a certain extent by varying the resonance frequency of thesecondary loop (by varying C3), the mutual coupling between the primaryand the secondary loop 32 and 34 is primarily dominated by the magneticcoupling and hence their proximity to one another. The range of mutualcoupling between the two loops by way of changing the resonance of thesecondary loop is small. Although, a greater range may be obtained byphysically moving the secondary loop, having moving coils in a RF coilpackage has not been a practical solution. Also, the closer thesecondary loop the greater will be the induced currents. Since thecurrents induced would be opposite in direction from the RF coilprimary, this will substantially reduce the net signal from the imagingFOV, which is not desirable.

Additional shortcomings associated with conventional RF coils havearisen due to limited control of the overlap between volume and surfacetype RF coils in the imaging FOV. Without the ability to control theextent of FOV overlap, particularly with regard to providing asymmetricoverlap, a reduced usable imaging FOV results in a reduced fillingfactor which causes a reduced S/N. Furthermore, oftentimes it isdesirable that an RF coil be tunable at two or more different NMRfrequencies. However, conventional RF coils capable of being tuned atdifferent frequencies either require traps for multiple tuning and/orresult in significantly different FOVs at the different frequencies.

In view of the aforementioned shortcomings associated with conventionalRF coil designs, there is a strong need in the art for an RF coil andmethod which provide the ability to vary the currents in the RF coil,such as to provide adequate RF screening, without significantlycompromising S/N in the imaging FOV. Moreover, there is a strong need inthe art for an RF coil arrangement and method which provides a high S/Nover the imaging FOV by allowing an asymmetric overlap between volumeand surface coils. In addition, there is a strong need in the art for anRF coil and method which provides for multiple tuning with similar Bfield profiles in the different NMR frequencies over the imaging FOV,without the use of traps.

SUMMARY OF THE INVENTION

The present invention provides an RF coil with a high S/N over theimaging FOV which is highly desirable for several MR studies (e.g.,anatomical, angiographic, functional, spectroscopic). The RF coil of thepresent invention exhibits adequate RF screening without significantlycompromising S/N over the imaging FOV. Moreover, the RF coil provides ahigh S/N and a high degree of B field homogeneity over the imaging FOV,by facilitating a unique combination of volume and surface coils. Inaddition, the RF coil may be multiply tuned for operation at two or moredifferent NMR frequencies while providing high S/N and similar B fieldprofiles over the imaging FOV.

In particular, the RF coil of the present invention provides the abilityto control the currents in the RF coil to provide efficient RFscreening, facilitates asymmetric overlap between surface and volumecoils in an array to obtain high S/N, and allows for simultaneousmultiple frequency operation. Specifically, the RF coil of the presentinvention provides high degree of RF screening without substantiallycompromising S/N over the imaging FOV; high S/N over the imaging FOV, byproviding asymmetric overlap between volume and surface coils; and highS/N and similar FOV's at the different NMR frequencies, by allowingcurrents at the different NMR frequencies to flow through the RF coilprimary. The RF coil exhibits a modified current distribution which isdifferent from the prior art.

The RF coil generally includes an RF coil primary, an RF coil secondary,and coupling impedances that connect the two at several points along thecoil peripheries. The coupling impedances are primarily reactive, andmay be fixed or variable, depending on the purpose of use.

The RF coil includes, but is not limited to, single or multi-channeloperation. The RF coil may be a surface or volume type coil and may besingly or multiply tuned for use in NMR or other resonance systems. Asurface coil design is best suited for breast, shoulder and torsoimaging, whereas a volume coil is best suited for whole body, head, kneeand wrist imaging, etc. Scaled down versions may be used to imagepediatric and premature neonates. Other applications that will benefitfrom the RF coil of the present invention are MR angiography,spectroscopy, functional and interventional MRI. The RF coil itself maybe used for transmit (T), receive (R), or transmit/receive (T/R)purposes, or may be used in conjunction with local gradients for veryhigh-resolution imaging.

According to one particular aspect of the invention, a radio-frequency(RF) coil for resonance imaging/analysis is provided. The RF coilincludes an RF coil primary sensitive to RF signals produced duringresonance imaging/analysis, the RF coil primary having a usable field ofview; an RF coil secondary positioned physically adjacent the RF coilprimary at a predetermined distance apart and mutually coupled to the RFcoil primary at a frequency of the RF signals; and coupling impedanceselectrically connecting the RF coil primary to the RF coil secondary toregulate an amount of current induced in the RF coil secondary at thefrequency, and to form current loops between the RF coil primary and theRF coil secondary serving to redirect at least a substantial portion offlux back into the useable field of view which would otherwise impingeon the RF coil secondary in the absence of the coupling impedances.

According to another aspect of the invention, a radio-frequency (RF)coil apparatus for resonance imaging/analysis is provided which includesan RF volume coil sensitive to RF signals produced during resonanceimaging/analysis; and an RF surface coil, also sensitive to the RFsignals, physically positioned relative to the RF volume coil to share acommon axis and to produce an overlap of the magnetic B fields of therespective coils at a frequency of the RF signals.

In accordance with still another aspect of the invention, a multipletuned RF coil is provided which includes an RF coil primary having ausable field of view; an RF coil secondary physically adjacent the RFcoil primary and mutually coupled to the RF coil primary; and couplingimpedances electrically connecting the RF coil primary to the RF coilsecondary to regulate an amount of current flowing between the RF coilprimary and the RF coil secondary, wherein the RF coil is tuned toresonate at a plurality of different RF signal frequencies producedduring resonance imaging/analysis.

According to still another aspect of the invention, an RF coil isprovided which includes an RF coil element having a surface covered witha precious metal selected from a group including gold and platinum.

To the accomplishment of the foregoing and related ends, the invention,then, comprises the features hereinafter fully described andparticularly pointed out in the claims. The following description andthe annexed drawings set forth in detail certain illustrativeembodiments of the invention. These embodiments are indicative, however,of but a few of the various ways in which the principles of theinvention may be employed. Other objects, advantages and novel featuresof the invention will become apparent from the following detaileddescription of the invention when considered in conjunction with thedrawings.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1a is a schematic diagram of a conventional loop-type receive onlyNMR coil;

FIG. 1b is a schematic diagram of a conventional decoupling network;

FIG. 2 is a schematic diagram of a conventional mutually coupled RFcoil;

FIG. 3 is a block diagram of an RF coil in accordance with the presentinvention;

FIG. 4a is a schematic diagram of a linear surface type RF coil inaccordance with a first embodiment of the present invention;

FIG. 4b is a schematic diagram of a secondary decoupling circuit for usein the RF coil of FIG. 4a;

FIG. 5 is a schematic diagram of a linear loop array type RF coil inaccordance with a second embodiment of the present invention;

FIG. 6 is a schematic diagram of a quadrature surface type RF coil inaccordance with a third embodiment of the present invention;

FIG. 7 represents a B field profile along the RF coil axis of FIG. 6;

FIG. 8a is a diagrammatic view of a system providing symmetric overlapbetween a quadrature volume coil and a quadrature surface coil;

FIG. 8b is a diagrammatic view of a system providing asymmetric overlapbetween a quadrature volume coil and a quadrature surface coil inaccordance with a fourth embodiment of the present invention;

FIG. 9 is a general schematic diagram of a quadrature, singly ormultiply tuned volume RF coil in accordance with a fifth embodiment ofthe present invention;

FIG. 10a is a detailed schematic diagram of an individual repeatingelement making up the RF coil of FIG. 9;

FIG. 10b is an equivalent model of the individual repeating element ofFIG. 10a at a low NMR frequency;

FIG. 10c is an equivalent model of the individual repeating element ofFIG. 10a at a high NMR frequency;

FIG. 11 is a graph illustrating the provision of adequate shieldingwithout compromising S/N in the imaging FOV in accordance with anexample of the prevent invention;

FIGS. 12a-12c are flux diagrams representing a conventional RF coil, anRF coil with conventional screening, and an RF coil in accordance withthe present invention, respectively;

FIG. 13 is a system diagram of a NMR imaging system incorporating an RFcoil in accordance with the present invention; and

FIG. 14a-14d are exemplary coupling impedances in accordance with thepresent invention.

DESCRIPTION OF THE PREFERRED EMBODIMENTS

The present invention will now be described with reference to thedrawings, wherein like reference numerals are used to refer to likeelements throughout.

The present invention relates to an RF coil and a method for modifyingthe magnetic B field profile for a RF coil to provide adequate RFscreening on one side of the RF coil without significantly compromisingS/N over the imaging FOV. The RF coil includes an RF coil primary and anRF coil secondary. The RF coil primary and secondary are coupled to eachother by mutual inductance and by one or more coupling impedances asdiscussed below in connection with FIG. 3. The B field of the RF coilmay be altered by making the currents in the RF coil secondary eitherinductive or capacitive, for example, so as to provide a symmetric or anasymmetric field. In doing so, the RF coil can also be multiply tuned.Should performance be critical at the low gamma nucleus, the reactiveimpedances such as the capacitors in the RF coil will be distributedsuch that most of the low frequency currents will circulate in the RFcoil primary with very little currents in the RF coil secondary. At thehigher NMR frequency or frequencies, the RF coil primary will behave asan RF short letting all the high frequency currents flow through. Such acoil is thus capable of simultaneous operation at multiple NMRfrequencies. The symmetric or asymmetric B field profile may also beused to facilitate a symmetric or asymmetric overlap between coils in anarray, and to provide a high S/N and a high degree of B fieldhomogeneity over the imaging FOV.

The RF coil of the present invention is described herein by way ofillustrative embodiments. Briefly summarizing the respectiveembodiments, the first embodiment is discussed below in connection withFIG. 4a and includes a planar loop type RF coil primary and a planarloop type RF coil secondary. The RF coil secondary is similar indimension and is disposed on one side to the RF coil primary so as toscreen one side of the RF coil primary, with the imaging FOV on itsother side. The RF coil primary and secondary are electrically coupledtogether by coupling impedances distributed about their peripheries.Such an RF coil represents a linear surface type coil.

The second embodiment of the present invention is a linear surface coilarray for uni or bi-lateral breast imaging as discussed below inconnection with FIG. 5. The array includes two solenoid type RF coilprimary loops surrounded respectively by two identical but largerdiameter solenoidal RF coil secondary loops, bridged by capacitancesdistributed around the coil peripheries. The capacitances may be fixedor variable to adjust the currents on the respective RF coilsecondaries, and may also be used for tuning the RF coils to the NMRfrequency.

The third embodiment is described in connection with FIG. 6 below. TheRF coil includes an eight leg, distributed capacitance-inductance typeRF coil primary, and an identical RF coil secondary placed adjacentalong a common coaxial axis connected by eight capacitors that bridgethe gap between the RF coil primary and the RF coil secondary along thecoil periphery. In this special case, the virtual ground points of theRF coil primary and secondary are tied together to serve as a commonground point. This coil represents a quadrature surface type coil.

According to another aspect of the invention, the surface type RF coilwhich is described in connection with FIG. 6 is combined with a volumetype coil. As is discussed below in relation to a fourth embodimentshown in FIG. 8b, the B field profile of the surface type RF coil isoverlapped in an asymmetric fashion along the coil axis with a birdcagetype volume coil. The net mutual coupling or mutual inductance betweenthe surface and volume coils is virtually zero. This helps to provideimproved S/N and B field homogeneity in the imaging FOV, e.g., overportions of the top of the head, when compared to a conventionalbirdcage type head coil.

The fifth embodiment is described below in relation to FIG. 9. The RFcoil is represented by a 16 leg, birdcage type RF coil primary, and a 16leg birdcage type RF coil secondary placed coaxially with a largerdiameter. The RF coil primary and secondary are connected by 32capacitors distributed along the end rings. Such embodiment represents aquadrature volume coil type. The RF coil may be a multiply tuned volumecoil capable of simultaneous operation at two or more differentfrequencies. For example, the coil may be tuned to operate at the ¹ Hand ³¹ P frequencies for MR spectroscopy.

Referring initially to FIG. 3, an RF coil in accordance with the presentinvention is designated generally as RF coil 40. The RF coil 40 includesan RF coil primary 42, an RF coil secondary 44, and coupling impedances46 that bridge the gap between the primary and secondary, physically andelectrically. The RF coil primary 42 and secondary 44 may be resonant atany given frequency, or in some cases be designed to be non-resonant.The coupling impedances 46 are mostly reactive elements. The couplingimpedances 46 may include an inductor, a capacitor, or a combinationthereof as discussed below in relation to FIGS. 14a through 14d. Therespective reactive elements can be fixed or variable and may be used tovary the currents in the RF coil primary 42 and/or secondary 44, to tunethe entire RF coil 40. Also, the currents in the RF coil secondary 44can be made either inductive (I_(si)), or capacitive (I_(sc)) dependingon the values of the coupling impedances 46. Hence, by adjusting thevalues of the coupling impedances 46 it is possible to achieve a desirednet effect on the performance of the RF coil primary 42 and hence the RFcoil 40, such as to produce a symmetric or an asymmetric B fieldprofile.

If the currents in the RF coil secondary 44 are inductive, then thesecondary currents _(si) will be in the opposite direction to that ofthe RF coil primary currents (I_(p)). Although, this will provide a neteffect on the performance of the coil 40, there will be RF screening.The extent of this screening including the net effect on the S/N willdepend on both I_(p) and I_(si).

Whereas if the currents in the RF coil secondary 44 are capacitive, thenthe secondary currents ISc will maintain the same direction as the RFcoil primary currents I_(p) and will provide a net effect on theperformance of the coil 40. Here, there may be little or no RFscreening. Depending on the current proportions in the RF coil primary42 and secondary 44, the result will be a symmetric or an asymmetric Bfield profile.

If the currents in the RF coil secondary 44 are nulled, that is, if thecurrents in the secondary 44 are neither inductive nor capacitive, thenthe RF coil secondary 44 will have no net effect on the performance ofthe RF coil primary 42 and hence the coil 40. In some cases this may benecessary where the RF coil 40 is tuned to one or more frequencies.

For a single tuned case of a whole body RF transmit coil, for example, alarge fraction of currents are induced on the secondary coil 44 whichare opposite in direction of the currents of the RF coil primary 42.This will substantially reduce the overall coil performance. Here, thecoupling impedances 46 can be varied such as to provide limited currentsin the RF coil secondary 44. Predetermined limited currents in the RFcoil secondary 44 will provide adequate screening without significantlyaffecting the performance of the RF coil primary 42 and hence the entireRF coil 40. This means the performance of the RF coil primary 42 willremain very close to that when operated without the RF coil secondarywith respect to the imaging field of view. This will drastically reducethe amount of RF power expended in the patient, with RF power depositionbeing significantly below the prescribed Food & Drug Administration(FDA) limits for specific absorption rate (SAR). Also, this willsignificantly reduce the amount of RF power expended in the RF coilsecondary 44 and hence the entire coil 40 during transmit.

As is explained in more detail below, the symmetric or asymmetric Bfield profile will facilitate a symmetric or an asymmetric overlapbetween coils in an array. This will assist in maintaining a high S/Nand a high level of B field homogeneity over the imaging FOV.

The coupling impedances 46 may also be used to tune the RF coil 40 to asingle or multiple NMR frequencies. To maintain a high S/N at the lowgamma nucleus (³¹ p, ¹³ C, ²³ Na, ¹⁷ O, etc.), it is preferred thatinitially the RF coil primary 42 be tuned close to the low NMR frequencyindependent of the RF coil secondary 44 (i.e., with the RF coilsecondary 44 open such that it has no net effect on the resonantfrequency of the RF coil primary 42). In such a system, at the lowfrequency the RF coil secondary 44 will behave like a open circuit. Atthe high frequency, the RF coil primary 42 will behave like a shortcircuit to the high frequency currents while allowing them to flowthrough. In such cases the imaging or the spectroscopic FOV are equallyspanned by the RF coil 40, since both the high and low frequencycurrents flow through the RF coil primary 42. As a result very similar Bfield profiles are created at the multiple NMR frequencies. This isextremely useful for obtaining ¹ H images, shimming, and for protondecoupling, all of which are important for obtaining spectroscopicinformation of the x-nucleus (³¹ p, ¹³ C, ²³ Na, etc.).

FIG. 4a illustrates a first preferred embodiment of an RF coil 40a. TheRF coil 40a includes an loop type RF coil primary 42a, resonating withC4, and a loop type RF coil secondary 44a. The coupling impedances46aare represented respectively by four C5 capacitors each of which maybe fixed or variable. Each of the RF coil primary 42a and secondary 44amay be, for example, loops measuring 12 cm×17 cm in size. The secondaryloop 44a may be resonant or non-resonant. In the illustrated embodimentcase, the secondary loop 44a is non-resonant. The RF coil primary 42aand the RF coil secondary 44a are separated by a distance D of 2.0 cm.The RF coil primary 42a is actively decoupled using the decouplingnetwork similar to that of FIG. 1b. The secondary 44a is made up fourlinear segments 1-14, each of which has a L2-C6-L3-D2 decoupling circuitI as represented in FIG. 4b. During transmit, the pin diode D2 isforward biased creating a trap and opening the segment, whereas duringreceive the L3-D2 circuit is open and the L2-C6 network is a shortcircuit at the NMR frequency. Typical values for L2-C6-L3 at 64 MHz for1.5T are 124 nH, 50 pF and 124 nH, respectively. The voltage across 2C4is matched to 50 ohms prior to the preamplifier. The entire coil systemis tuned to 64 MHz, which happens to be the ¹ H frequency for a 1.5Tesla B_(o) field strength.

The currents in the RF coil secondary 44a may be altered by varying theC5 capacitors, which will also de-tune the coil 40a. The coil may thenbe fine tuned by using trimmer capacitors (not shown) across C4 or C5.Depending on the size of the RF coil 40a, the separation distance Dbetween the primary 42aand secondary 44a, and the operating frequency,capacitors C4 or C5 may have greater control in the coil tuning. Forthis specific embodiment, C5 has been found to have greater control overthe coil tuning. In other words, a slight variation in C5 resulted in awide de-tuning of the entire coil 40a. However, the slight re-tuning ofthe coil 40a did not markedly alter the currents in the RF coilsecondary 44a. Therefore, by fine tuning using a small value trimmercapacitor, adequate RF screening and precise tuning can be achieved. Inall cases, the RF coil primary 42a and secondary 44a assume fixedpositions, that is the RF coil primary and the RF coil secondary neednot be physically moved to achieve the desired RF screening and tuning.

Should the coupling impedances 46a present a RF short between the RFcoil primary 42a and secondary 44a, a larger fraction of currents willflow in the secondary and in this special case, the secondary currentswill be inductive and in the opposite direction of the RF coil primarycurrents. However, should the coupling impedances 46a present an opencircuit, the fraction of the currents in the RF coil secondary 44awilldepend on the physical distance D between the primary 42a and thesecondary 44a and on the resonance of the secondary. Since the distanceis fixed in this case, a fixed relatively small fractional current willbe present in the secondary 44a. This can be altered dramatically usingthe coupling impedances 46a mentioned above.

With such control over a large range without moving the secondary, thecurrents can be altered so as to maintain an adequate screening. Giventhe RF coil secondary 44a is fixed in location and placed at a suitabledistance from the RF coil primary 42a, the effective RF screening can beachieved without significantly compromising the S/N over the imagingFOV. This is because the electrically forced currents (via adjustment ofthe coupling impedances 46a) in the secondary 44a (inductive in thiscase) have a minimal detrimental effect over the imaging FOV. The aboveis an example of a self-screening linear surface coil. It is noted thatadditional fine tuning can be carried out by adjusting one or more ofthe other sets of capacitors and inductors within the coil 40a.

FIG. 5 is a second preferred embodiment and is designed for uni- orbi-lateral breast imaging in an MRI system. Here two such circular loopsolenoid type RF coil primaries 42b with capacitors C6 are connected totheir respective non-resonant solenoid type RF coil secondaries 44b viacoupling impedances 46b in the form of capacitors C7. It is noted thatthe RF coil secondaries 44b may be resonant as well, where a combinationof the impedances in the coupling elements including the reactiveimpedances in the secondary will alter the currents in the RF coil. Inthis embodiment, the RF coil primaries 42b each have similar decouplingnetworks 22 of FIG. 1b. The secondaries 44b are equipped with similardecoupling networks of FIG. 4b (not shown). The currents in the RF coilsecondaries 44b can be controlled to provide adequate screening betweencoils, between the coils and near by anatomy, and still provide highenough S/N in the imaging volume for reasons similar to those describedabove in connection with the embodiment of FIG. 4a. The left and rightcoils 40b can be turned ON or OFF selectively using the DC control linesfrom an MRI system, such that the coils are used to image in auni-lateral or a bi-lateral fashion. Individual coils 40b are matched to50 ohms and fed to the individual preamplifiers. This is an example of asurface coil array.

Turning now to FIG. 6, a quadrature surface coil 40c is shown inaccordance with the third embodiment of the present invention. Theprimary RF coil 42c is a circular loop connected by eight straightsegments to a virtual ground point VG1 at the primary coil center. Eachof the eight segments are split by two capacitors C8, whereas thecircular loop is split by four high value RF short capacitors C9. One ofthe two C8 capacitors is decoupled using a decoupling network (notshown) similar to that of FIG. 1b. The circular loop is split by verylargesplit by very large value capacitors C9 to reduce any gradiencurrents, while letting RF currents flow. The RF coil secondary 44c isidentical in dimension and is placed adjacent to the RF coil primary42c, and maintains a common coil axis. The RF coil secondary 44cincludes eight straight segments similar in orientation to the segmentsin the RF coil primary 42c. Each of the eight segments of the RF coilsecondary 44c have L-C-L-D decoupling circuits similar to that of FIG.4b. The loop of the RF coil secondary 44c is broken similar to the RFcoil primary 42c, with high value RF shorting capacitors C10(C10=C9=1000 picofarads (pF)). Both the RF coil primary 42c andsecondary 44c maintain their own virtual ground points, VG1 and VG2,respectively. Eight C11 capacitors serve as the coupling impedances 46cwhich bridge the gap between the RF coil primary 42c and the secondary44c and are azimuthally distributed along the coil periphery. The entirecoil 40c has multiple frequency modes, and the principal quadrature modeis resonant at the NMR frequency.

The principle linear modes of this quadrature coil are matched to 50ohms, prior to the preamplifier (not shown). The coupling ports A and Bare shown in FIG. 6. The center conductor of these two coaxial cablesare connected to capacitors across ports A and B, via reactive means tomatch the individual ports to 50 ohms. As seen, one of the points ofthese coupling ports A and B is forced to ground. Also, coaxial cables(not shown) exiting the RF coil primary at VG1 are tied to VG2, beforeexiting to the MRI system. This forces VG1 and VG2 to the system groundvia the shields of the coaxial cables, which eliminates the need forcable traps and such mechanisms commonly used to reduce currents on thecoaxial cable shields.

The currents in the RF coil secondary 44c may be altered similar to theabove embodiments, that is by varying the coupling impedances 46c (C11).Also, the entire coil 40c may be fine tuned by doing the same. The twolinear ports A and B are then combined using analog combiners prior tothe preamplifier. These two ports may also be directly interfaced to twodifferent channels of a multichannel system via individualpreamplifiers. The currents in the RF coil primary 42c segments follow acosinusoidal distribution. Since a fraction of these currents exist inthe secondary 44c, and since all C11 values remain very close, thecurrents in the RF coil secondary 44c also follow a cosinusoidaldistribution. For a detailed explanation of the theory of the RF coilprimary design, please refer to the article by Srinivasan, et al,published in the 4th Scientific Meeting of the International Society ofMagnetic Resonance in Medicine, Book of Abstracts, New York, Apr. 27-May3, 1996, page 1425.

Inductive currents in the RF coil secondary 44c can assist inaccomplishing RF screening, where the RF coil 40c will maintain anasymmetric B field profile from one side of the coil 40c to another.FIG. 7 is a normalized B field simulation along the coil axis fordifferent secondary currents for the design of FIG. 6. Coil dimensionsare 22.5 cm in diameter, 8 legs and a primary 42c and secondary 44cseparation distance D of 2 cm. As seen for I_(si) =-0.4I_(p), the Bfield profile is highly asymmetric and the screening is greater than forI_(si) =0. The negative sign indicates that the inductive currents onthe secondary are opposite to that of the primary I_(P). Please note,I_(si) =0 is for a unscreened case for this specific coil design.

FIG. 8a illustrates an overlap between a volume coil 60 and a surfacecoil 62 which maintain a common axis. The surface coil 62 is of thequadrature type similar to the RF coil primary 42c of FIG. 6, whereasthe volume coil 60 is of the quadrature birdcage type most commonly usedfor head imaging. Since the surface coil 62 is un-screened, this coil 62will exhibit the same B field profile along the coil axis and on boththe coil sides. Therefore, a suitable place where the cross-talk betweenthe volume coil 60 and the surface coil 62 is kept to minimum is at thecentral virtual ground plane 64 of the volume coil 60. At this plane 64,the mutual coupling is virtually zero. That is, the flux coupling fromone coil to the other is almost zero, and the coils 60 and 62 maintainminimum mutual inductance.

As seen in FIG. 8a the coil array does not allow for placement of theentire head inside the volume coil 60. Also, the filling factor ismarkedly reduced which results in a reduced S/N. Accordingly, it isdesirable that there be a coil design in which the surface coil can beplaced off center and along the coil axis, such that the entire head canbe placed inside the volume coil. In order to achieve this goal, the Bfield profile of the surface coil 60 has to be modified, that is the Bfield profile has to be asymmetric. This means that the RF field on oneside of the surface coil and along the coil axis had to be screened to acertain extent. The asymmetric overlap can now be achieved with asurface coil that maintains an asymmetric B field profile along the coilaxis as shown in FIG. 8b.

Specifically, the inventor has overlapped a surface coil 42c of the typedescribed with respect to FIG. 6 with the birdcage type high-pass volumecoil 60. The volume coil 60 has a 30 cm diameter and is 30 cm in lengthwith sixteen legs connecting the two end rings. The surface coil 40c isa 22.5 cm diameter, 8 segment RF coil as represented in FIG. 6, having a2.0 cm separation between the RF coil primary 42c and secondary 44c. Thesurface coil 40c, tuned for an asymmetric B field, can be placed towithin 5 cm from one end ring of the birdcage volume coil 60. Thisallows an entire head to be placed inside the volume head coil from theother end of the coil 60, while still maintaining a high S/N and a highdegree of B field homogeneity along all three planes of the coil. Morespecifically, a high S/N is obtained toward the top of the head idealfor MRI applications such as functional magnetic resonance imaging MRIand MR angiography where the focus is from the circle of willis to thetop of the head.

Thus the asymmetric B field profile of the quadrature surface coil 40cfacilitates an asymmetric overlap between volume and surface coils thatmaintain a common axis. The asymmetric profile mandates their degree ofoverlap. The greater the asymmetricity of the B field along the surfacecoil 40c axis, the smaller the overlap between the birdcage and thesurface coil, and vice versa is also true. That is, the greaterasymmetricity the smaller is the overlap where a minimal mutualinductance is easily attained, between the volume and surface coils inthe array. Minimum mutual inductance will result in reduced noisecorrelation between coils in the array, and therefore will result inincreased S/N in the combined image.

The quadrature birdcage coil 60 is interfaced to one channel and thequadrature surface coil 40c is interfaced to one other channel of a 1.5Tmultichannel MRI system (not shown). Alternatively, the birdcage coil 60linear modes could be interfaced to two channels and the two linearmodes of the surface coil could be interfaced to two other channels ofthe MRI system. FIG. 9 represents a fifth embodiment of the presentinvention. The RF coil 40d includes a primary RF coil 42d which is a 16leg, birdcage volume coil of the low-pass configuration. Each of the 16legs is broken by a C12 capacitor as shown more clearly in FIG. 10a. TheRF coil 40d further includes an RF coil secondary 44d which is also abirdcage type volume coil, except has a larger diameter than the RF coilprimary 42d while having the same length. The legs of the RF coilsecondary 44d have no capacitors. The RF coil primary 42d and secondary44d are bridged by 32 capacitors C13 serving as coupling impedances 46d,azimuthally distributed on both sides along the coil periphery. Notshown are capacitors that split the end rings of the RF coil primary 42dand secondary 44d, to reduce the gradient induced eddy currents. Thecoil 40d exhibits a multitude of frequency modes, but only two frequencyquadrature modes are observed at the coil center. These are the twofrequency modes where the coil is tuned and operated, simultaneously.

A coil 40d with a RF coil primary 42d of 18 cm diameter and a RF coilsecondary 44d of 24 cm diameter was built by the inventor. The RF coilprimary 42d and secondary 44d were each 19 cm long and had 16 legs each.The RF coil primary 42d was of the low-pass configuration. The RF coilsecondary 44d was non-resonant. The capacitor C12 values were 68 pF, andthe capacitor C13 values were 92 pF. The coil 40d was doubly tuned to24.1 MHZ and 65.3 MHZ. These frequencies were close to 64 and 26 MHZ,for ¹ H and ³¹ p at 1.5T. Therefore, the coil 40d could then be finetuned with small additional capacitances, to tune to the respective NMRfrequencies.

An individual repeating element of the coil of FIG. 9 is shown in FIG.10a. It is noted that since this circuitry has five loops (one in the RFcoil primary, one in the RF coil secondary and three joining both),eigen-value solutions for estimating the several frequency modes ω_(k)will be a fifth order equation. The individual repeating elements atboth the low and high NMR frequencies are reflected in FIGS. 10b and10c, respectively.

As represented from FIG. 10b, at 26 MHZ most of the currents circulatein the RF coil primary 42d, and the RF coil secondary 44d appearedsubstantially as an open circuit. A very small fraction of currentcirculated in the secondary 44d. This was confirmed by isolating the RFcoil primary 42d from the coil system and the coil resonating at 25.7MHZ. Since most of the low frequency currents circulate in the RF coilprimary 42d, the efficiency at the low frequency will be similar to thesingle tune design. The low frequency coil may thus be modeled as shownin FIG. 10b.

At 64 MHZ, the RF coil primary 42d presents a short circuit letting allthe high frequency currents to flow. Thus the RF coil 40d can be modeledsimilar to FIG. 10c, at the high NMR frequency.

Since the RF coil primary 42d carries a major fraction of both the highand low frequency currents, the RF coil 40d inventive design willmaintain a similar FOV at both the NMR frequencies. This is essentialfor MR spectroscopy of ¹ H or other nuclei referred to above as thex-nucleus. Also, the sensitivity of the high NMR frequency will remainclose to the single tune design. Thus, this design will maintain highefficiencies at both the NMR frequencies. Also, this coil design willallow simultaneous operation, which is essential for proton-decouplingand cross-polarization experiments.

It is noted that although the preferred embodiment has 16 legs in the RFcoil secondary 44d, in another embodiment the coil may include adifferent number of legs, e.g., 8 legs. In yet another embodiment, theRF coil secondary 44d may be replaced with a solid or fine mesh RFscreen. In either of such cases, the invention will still be resonant inquadrature at the two NMR frequencies. It is noted that a quadraturemode has two linear modes, which in turn can be tuned to differentfrequencies. Therefore the quadrature, double tuned birdcage design ofFIG. 9 can be tuned and operated at four different NMR frequencies,simultaneously.

It is also noted that the coil 40d of FIG. 9 can be operated in a singletuned mode with RF screening. Such coil will find ready application forknee imaging, where the contralateral asymptomatic knee is close to thecoil. A self-screened RF coil will allow for a free direction offrequency and phase encoding in NMR and will help reduce chemical shiftand blood flow artifacts. This will improve the quality of the MR image,which in turn will assist the clinical diagnosis.

FIG. 11 is a B field profile illustrating an example of the effect thepresent invention has with regard to providing RF screening withoutsacrificing S/N over the FOV. Presume a prior art RF coil secondary ispositioned relative to an RF coil primary 42 in order to provideequivalent screening (B field) as the RF coil secondary 44 with couplingimpedances 46 as represented by curve 60. In each case the currents inthe RF coil secondary are opposite in direction to the currents in theRF coil primary to provide RF screening. The inductive currents in theRF coil secondary for the prior art will be greater than the inductivecurrents of the RF coil secondary in the present invention. This willreduce the net effect on the S/N over the imaging FOV. The RF profileover the imaging FOV for the present invention is given by curve 62. TheRF profile over the imaging FOV for the prior art arrangement is givenby curve 64. Hence, given the same RF screening the present inventionprovides improved S/N over the imaging FOV.

FIG. 12a presents a flux diagram for an unscreened RF coil 20 such asthe type discussed above in connection with FIG 1a. As noted, the fluxfield lines B are identical on both sides of the coil 20. Here, there isno RF screening.

In comparison, FIG. 12b illustrates a flux diagram for an RF coil 30such as the type shown in FIG. 2. As previously discussed, the RF coilprimary 32 is mutually coupled through space to the RF coil secondary34. RF currents will be induced in the secondary 34, which according toLenz's Law, will oppose the very currents in the primary 32 whichproduced them. That is, the RF currents on the secondary 34 will be inthe opposite direction to that of the primary 32; their proportion,however will be mandated by their proximity (e.g., distance D). In thecase shown in FIG. 12b, the majority of the impinging flux B willterminate at the secondary 34. However, a small fraction will bediverted to the usable imaging coil FOV. The extent of the fluxtermination will also depend on the proximity of the secondary 34 to theprimary 32. Although there will be screening on one side of the coil 30,the net flux in the usable imaging FOV will be reduced from theun-screened case of FIG. 12a. This will result in a reduced S/N in theimaging FOV.

FIG 12c exemplifies the utility of the present invention. FIG. 12crepresents the flux diagram for an RF coil in accordance with thepresent invention such as the type shown in FIG. 4a above. In this case,the majority of the flux B is compressed between the primary 42a and thesecondary 44a and is forced back into the usable imaging FOV. This isdue to the currents circulating in the loops established by the couplingimpedances 46a connecting the primary 42a and the secondary 44a. Thesecurrents maintain a direction such as to have their resultant B fieldsin the outward normal direction, which serve to force the flux B in suchdirection thus diverting their resultant B fields in to the usable FOV.A minority of flux will impinge on the secondary 44a, so as to providescreening. However, the amount of flux which impinges on the secondary44a is substantially less than the amount of flux impinging on thesecondary 34 in FIG. 12b when the respective primaries and secondariesare both a distance D apart.

Thus, according to the present invention by selecting the appropriatecoupling impedances 46 between the primary and secondary it is possibleto a) regulate the amount of induced currents in the secondary toprovide adequate screening; and b) divert the remaining flux back intothe usable imaging FOV. Consequently, it is possible to maintain a highS/N when compared to the prior art of FIG. 12b, for example, and closeto the same S/N in the unscreened case as in FIG. 12a. It is noted thatthe S/N for the invention will be slightly lower than the un-screenedcase due to loss of flux in the secondary to provide RF screening.However, the combination of RF screening and S/N is superior than ineither conventional coil.

The system block diagram of FIG. 13 illustrates the utility of the RFcoil 40 of the present invention in NMR imaging, for example. The systemhas a main magnet 70 which covers the time varying gradient coils, an RFshield 72 that shields the RF coil from the gradient coils 73 and awhole-body RF coil 74 most commonly used for uniform B field transmitover a large imaging FOV. The main magnet field strength sets the NMRfrequency. The time varying magnetic fields help spatially encode theNMR signals. The RF whole body coil 74 is used to transmit, while thelocal RF coil 40 is used to pick up the NMR signals from the objectunder investigation (NMR phantom 75). A number of receiver coils may beused in an array configuration in the RF coil 40 and may be summedeither analog or digitally to produce the resultant image. Signals fromthe i several receiver ports may be acquired via one or multiplereceiver channels. An n to 1 channel multiplexer is shown in thedrawing. This helps by-pass n channel coil data to use one channel ofthe NMR system. Alternatively, an n channel NMR system may also be used.

FIGS. 14a through 14d illustrate different examples of the types ofcircuits which can be used as the coupling impedances 46. The variousparallel/series relationships of the reactive elements can be selectedbased on the desired properties. One or more of the elements can bevariable to allow for tuning as desired.

It will be appreciated based on the proceeding description that the netor resultant current in the RF coil secondary 44 can be inductive,capacitive or neither (i.e., nulled). The B field of the RF coil 40 canbe symmetric or asymmetric along the axis of the coil. Volume andsurface coils can be symmetrically or asymmetrically overlapped tomaintain minimum mutual inductance.

Although the invention has been shown and described with respect tocertain preferred embodiments, it is obvious that equivalents andmodifications will occur to others skilled in the art upon the readingand understanding of the specification. For example, after reading theabove disclosure it will be apparent to someone having ordinary skill inthe art that the ability to vary the current in the RF coil primary andsecondary can be applied to almost all surface and volume type coils foruse in NMR. Examples of the linear and quadrature coil types are looptype, FIG. 8 type, solenoidal type, saddle type, distributedinductance-capacitance type, birdcage type, loop gap resonator type,etc.

It is also noted that there may be more than one RF coil secondary in adesign, the multiple secondaries being coupled to the RF coil primaryvia a common or independent set of coupling impedances. Furthermore, itis noted that the coil designs may be of the low-pass, high-pass, orband-pass variety or their combination. Also, the coils can be singlytuned or multiply tuned or may be used independently or in an arrayconfiguration. Principal k=1 modes or higher order k>1 modes may betuned to the NMR frequency. Inductive magnetic coupling or electricalcoupling may be used to match the individual modes of the coil to 50ohms.

The coils may also be used for transmit only, for receive only, or forboth transmit and receive of NMR signals. Furthermore, the RF coil canbe used for other techniques such as electron spin resonance, nuclearquadruple resonance, etc. The coils may be used alone or in conjunctionwith local gradient set up for high resolution and rapid imaging.

Also, the shape and form of the RF coil can be adapted to the surface orvolume under investigation. The RF coil primary and the RF coilsecondary may be resonant or non-resonant. The RF coil secondary alsomay be a non-resonant solid or fine mesh RF screen. The RF coilsecondary may be of a different shape and design than the RF coilprimary. The impedance coupling network may be reactive or resistive, ora combination of both.

Preferably the RF coil primary and secondary are made of copper, eitheretched to a rigid or a flexible printed circuit board, or used as amilled copper sheet of definite thickness. Copper tubes may also beemployed depending on the use of purpose including the product packagingneeds. For example, the whole body RF coil is generally made of coppertubes, whereas the surface coil is normally etched on a flexible printedcircuit board. It is noted that the RF coils may be covered with aprecious metal to enhance their performance, and to minimize oreliminate the oxidation of copper over time. Tin is used to reduce oreliminate the oxidation of copper, however tin does not enhance the coilperformance. The RF coils may be coated with precious metals such assilver, gold or platinum, etc., which have increased conductivity, tohelp reduce the resistance and increase the Q factor of a RF coil(Q=ωL/R). Since S/N is proportional to the square root of the coil Q,any increase in Q will result in an increase in the coil S/N and willresult in a lower transmitter power needed for the same experiment.

The degrees of freedom in coil design are its shape, including thediameter, coil length, number of legs and the separation between theprimary and secondary. An additional degree of freedom in coil design isthe ability to control the currents in the RF coil primary andsecondary, such as to provide adequate RF screening withoutsignificantly compromising S/N over the imaging FOV. This will alsofacilitate asymmetric overlap between volume and surface coils includingmultiple tuning of this novel structure.

The present invention includes all such equivalents and modifications,and is limited only by the scope of the following claims.

What is claimed is:
 1. A radio-frequency (RF) coil for resonanceimaging/analysis, comprising:an RF coil primary sensitive to RF signalsproduced during resonance imaging/analysis, the RF coil primary having ausable field of view; an RF coil secondary positioned physicallyadjacent the RF coil primary at a predetermined distance apart andmutually coupled to the RF coil primary at a frequency of the RFsignals; and coupling impedances electrically connecting the RF coilprimary to the RF coil secondary to regulate an amount of currentinduced in the RF coil secondary at the frequency to provide screeningrelative to the RF coil primary, and to form current loops between theRF coil primary and the RF coil secondary serving to redirect at least asubstantial portion of flux back into the useable field of view whichwould otherwise impinge on the RF coil secondary in the absence of thecoupling impedances.
 2. The RF coil of claim 1, wherein the RF primaryand secondary coils each comprise a planar loop type coil.
 3. The RFcoil of claim 2, wherein the loop type coils of the RF primary andsecondary coils are respectively oriented in parallel planes.
 4. The RFcoil of claim 3, wherein the loop type coils are rectangular.
 5. The RFcoil of claim 3, wherein the loop type coils are circular.
 6. The RFcoil of claim 3, wherein the RF coil is a quadrature coil.
 7. The RFcoil of claim 1, wherein the RF coil secondary is resonant.
 8. The RFcoil of claim 1, wherein the RF coil primary comprises a planar coilhaving an axis substantially perpendicular to the plane thereof, and amagnetic B field profile of the RF coil at the frequency is asymmetricalong the axis.
 9. The RF coil of claim 1, wherein the RF primary andsecondary coils each comprise a volume type coil.
 10. The RF coil ofclaim 9, wherein the volume type coils are each birdcage type coils. 11.A resonance imaging/analysis system, comprising;an RF coil as recited inclaim 1; and means for processing RF signals which are at least one ofreceived from the RF coil and transmitted from the RF coil in order toobtain a resonance image/analysis.
 12. A multiple tuned RF coil,comprising:an RF coil primary having a usable field of view; an RF coilsecondary physically adjacent the RF coil primary and mutually coupledto the RF coil primary; and coupling impedances electrically connectingthe RF coil primary to the RF coil secondary to regulate an amount ofcurrent flowing between the RF coil primary and the RF coil secondary,wherein the RF coil is tuned to resonate at a plurality of different RFsignal frequencies produced during resonance imaging/analysis.
 13. TheRF coil of claim 162 wherein a B field profile in the usable field ofview is substantially similar at the plurality of different RF signalfrequencies.
 14. The RF coil of claim 12, wherein the RF coil primaryrepresents substantially an RF short circuit at one of the plurality ofdifferent frequencies and the RF coil secondary represents substantiallyan RF open circuit at another of the plurality of different frequencies.15. The RF coil of claim 12, wherein the RF coil secondary isnon-resonant.
 16. The RF coil of claim 12, wherein the RF coil secondaryfunctions as an RF screen.
 17. A radio-frequency (RF) coil for resonanceimaging/analysis, comprising:an RF coil primary sensitive to RF signalsproduced during resonance imaging/analysis, the RF coil primary having ausable field of view; an RF coil secondary positioned physicallyadjacent the RF coil primary at a predetermined distance apart andmutually coupled to the RF coil primary at a frequency of the RFsignals; and coupling impedances electrically connecting the RF coilprimary to the RF coil secondary to regulate an amount of currentinduced in the RF coil secondary at the frequency, and to form currentloops between the RF coil primary and the RF coil secondary serving toredirect at least a substantial portion of flux back into the useablefield of view which would otherwise impinge on the RF coil secondary inthe absence of the coupling impedances.